Chronic performance control system for rotodynamic blood pumps

ABSTRACT

In a left ventricular assist device (LVAD) a rotodynamic blood pump ( 10 ) is powered by a brushless DC motor ( 12 ). A power supply ( 14 ) supplies power to the motor ( 12 ). Three feedback channels, one for each of voltage, current, and motor speed lead to a microcontroller or microprocessor ( 18 ). The three feedback waveforms are analyzed, and from these waveforms, motor input power, patient heart rate, current pump flow rate, and systemic pressure are determined. The microprocessor ( 18 ) then calculates a desired flow rate proportional to the patient heart rate. The microprocessor communicates a new power output to a commutation circuit ( 16 ), which regulates power to the motor ( 12 ). The pump ( 10 ) also includes safety checks that are prioritized over desired pump flow. These include prevention of ventricular suction, low pulsatility, minimum and maximum pump speed, minimum speed-relative pump flow, minimum absolute pump flow, minimum and maximum motor input power.

CROSS REFERENCE TO RELATED APPLICATION

This is a U.S. National Phase Application of PCT/US01/08776, filed Mar.19, 2001, which claims the benefit of U.S. Provisional Application No.60/192,221, filed Mar. 27, 2000.

FEDERAL RESEARCH STATEMENT

The U.S. Government may have certain rights in this invention pursuantto contract number N01-HV-58159 awarded by the U.S. National Heart, Lungand Blood Institute of the National Institutes of Health

BACKGROUND OF THE INVENTION

The present invention relates to the medical arts. It finds particularapplication in cardiac assist technologies using, for example,rotodynamic blood pumps, also known as left ventricular assist devices(LVAD) in assisting patients with failing hearts and will be describedwith particular reference to a centrifugal blood pump. It is to beappreciated that the present invention is also applicable to other typesof pumps, such as axial flow pumps, and is not limited to theaforementioned application.

Electrically driven rotodynamic pumps (axial flow, mixed flow andcentrifugal) have prospective applications in cardiac assisttechnologies. A typical cardiac assist system includes the blood pumpitself, electrical motor (usually a brushless DC motor integrated intothe pump), drive electronics, microprocessor control unit, and an energysource, such as rechargeable batteries. These pumps are used in fullyimplantable systems for chronic cardiac support. In this case the wholesystem is located inside the body and there are no drive linespenetrating the skin. For temporary support, and as well as for thebridge-to-transplant application, the pump itself is also located insidethe body. However some system components including drive electronics andenergy source may be placed outside the patient body.

Both chronic and temporary patient support require controlling the pumpperformance to satisfy the physiologic needs of the patient whilemaintaining safe and reliable system operation.

The primary goal for cardiac assist control is to provide an adequateblood pump flow rate for the patient that may depend on variousphysiological and psychological factors. Prior systems have pressuresensors, or ECG sensors external to the pump to determine the heart rateand blood pressure of the patient. These systems require extra hardwareinside the patient and increase the risk of complication.

U.S. Pat. No. 5,888,242 to Antaki et al. discloses a rotodynamicventricular assist pump that uses current measurements and pumprotations per minute (rpm) measurements to test and identify a maximumrpm that will not cause ventricular collapse. The invention described inthis patent monitors for a current spike indicative of ventricularcollapse, and in response decreases the pump speed. It does thisiteratively to achieve a maximum average flow rate. This approach putsunnecessary strain on the heart by continuously depending on thisdangerous situation to optimize pump flow.

The present invention provides a new and improved method and apparatusthat overcomes the above referenced problems and others.

SUMMARY OF THE INVENTION

In accordance with one aspect of the present invention, a cardiac assistdevice is provided. A blood pump is driven by a drive unit powered by apower supply. Three measurable parameters, current, voltage androtational frequency (or pump speed) each relay their respective sensedwaveforms to a controller. A motor winding commutation circuit directedby the controller directs power to the drive unit in response to thesensed waveform.

In accordance with another aspect of the present invention, a method ofcontrolling blood flow with a blood pump is given. A current waveform,voltage waveform, and rotational frequency waveform, are sensed. Thesensed information is provided to a blood pump controller which altersoperation of the blood pump if necessary.

One advantage of the present invention is its independence of supplyvoltage variations.

Another advantage is that it uses a sensorless approach where no flow orpressure sensors are used.

Another advantage is simple control circuitry.

Another advantage is that it takes into account blood viscosityvariations.

Yet another advantage resides in the enhanced control flexibility andprecision while improving system reliability and effectiveness.

Still further benefits and advantages of the present invention willbecome apparent to those skilled in the art upon a reading andunderstanding of the preferred embodiments.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention may take form in various components and arrangements ofcomponents, and in various steps and arrangements of steps. The drawingsare only for purposes of illustrating preferred embodiments and are notto be construed as limiting the invention.

FIG. 1 is a graphical representation of targeted pump flow (Q) to heartrate (HR).

FIG. 2 is a graphical representation of targeted pump flow (L/min)versus a ratio (M) of heart rate (HR) and mean systemic pressure (P).

FIG. 3 is a graphical illustration of a normalized pump pressure-flowcurve.

FIG. 4 is a graphical representation of normalized power plotted againstflow.

FIG. 5 is an illustration of pulse width modulation used to controlmotor excitation power.

FIG. 6 is a diagrammatic illustration of a left ventricular assistdevice drive unit in accordance with the present invention.

FIG. 7 is a graphical representation of blood viscosity versushematocrit level.

FIG. 8 is a flow diagram of safety checks that prioritizes emergencyconditions.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

There are several possible approaches to setting a required pumpperformance level matched to patient individual needs. A simple oneemploys a few levels of fixed pump speed/flows (i.e., low, medium, high)that are set according to patient activity (i.e., sleeping, sitting,walking). Another more flexible approach uses cardiac pacing technologywhere a separate pacemaker type device controls the pump performance.The present invention, though, proposes methods that employ patientsystemic pressure and/or heart rate to define the required pump flowwhile using no pressure or ECG type sensors (so-called “sensorless”approach).

Studies have shown that there is an almost linear dependence betweenhealthy human heart rate and cardiac output for different age groups andvarious activity levels. Therefore, maintaining an appropriate pump flow(Q_(target)) depending on the patient heart rate (HR) is a reasonableobjective for the heart assist control algorithm.

It is important that the controlled pump flow response to changes inpatient HR be limited by a maximum (Qmax) and minimum (Qmin) allowedpump flow associated with upper (HRmax) and lower (HRmin) heart ratelimits respectively. In other words, if HR>HRmax then Q_(target)=Qmax;if HR<Rmin then Q_(target)=Qmin. The values of Qmax, Qmin, HRmax andHRmin are based on the patient's residual ventricular function, bodysize and activity level. In the preferred embodiment (see FIG. 1.) thepump flow demand is proportional to the heart rate:Q _(target) =C ₁ *HR+C ₂, L/min.  (1)where the constants C₁ and C₂ are found using following expressions:C ₁=(Q _(max) −Q _(min))/(HR _(max) −HR _(min))C ₂ =Q _(max) −C ₁ *HR _(max).

A linear dependence between the patient heart rate and the assistingpump flow rate is not the only possible relationship. There is somemerit to an argument that the pump flow response can be a non-linearfunction. Generally, any equation that defines the required pump flow asa function of the patient heart rate can be used.

An alternate approach uses the patient heart rate and an estimated meansystemic pressure P to control the pump flow (see FIG. 2). The ratioM=HR/P is used as an independent variable that defines the appropriatepump flow:Q _(target) =f(M).  (2)where f is a monotonic function within the specified interval [Mmin,Mmax]. It is also important that the targeted pump flow response tochanges in the M ratio be limited by maximum and minimum M values thatare defined as follows:Mmax=HRmax/Pmin;Mmin=HRmin/Pmax.Where Pmax and Pmin are upper and lower pressure limits respectively. Inother words, if M>Mmax then Q_(target)=Qmax; if M<Mmin thenQ_(target)=Qmin. Therefore, by setting Mmax and Mmin, a maximum andminimum pump flow response is defined. Again, the values of Qmax, Qmin,HRmax, HRmin as well as Pmax and Pmin are defined by the patient'sresidual ventricular function, body size and activity level. It is to beunderstood that any function that is monotonic within specified interval[M_(min), M_(max)] can potentially be employed to define the requiredpump flow.

The relationship between pump flow and motor input power over a definedpump operating range can be obtained from the steady state power-flowcharacteristic for a particular pump type. Neglecting fluid inertia thisrelationship can be used to obtain instantaneous pump flow from a motorpower waveform. Detection of motor power is both a safe and reliablecontrol feedback for motors allowing determination of the instantaneousblood pump flow without direct measurement of flow or pressure dropacross the pump. The detected instantaneous flow waveform can beemployed to find the pump mean flow. The comparison of the target meanflow value derived from physiologic HR or HR/P input and the currentcalculated pump mean flow, results in an error signal used to change thepump performance in order to achieve the mean (target) blood flow.

When the rotodynamic blood pump is operated as an LVAD with ventricularcannulation, the motor power, current, speed and pump flow pulsatilitycorrespond with the pulsatility of the ventricular pressure input intothe pump inflow cannula. Thus, the heart rate may be found by analyzingthe frequency components of the motor current, or speed, or power, orflow waveforms. The fundamental frequency of all of these waveforms isdefined by the heart beat rate. This allows heart beat rate detectionwithout direct ECG sensing.

The speed waveform and the calculated pump flow waveform then are usedto calculate the pump pressure drop waveform from which the maximumpressure drop across the pump is derived. The maximum pressure dropacross the pump is used to estimate mean systemic pressure P. Anormalized pump pressure-flow curve is used that is independent of pumpoperational speed:P _(norm)=ψ(Q _(norm))  (3)where normalized pump pressure P_(norm)=P/(σ*ω²), normalized flow rateQ_(norm)=Q/ω, σ is the blood density and ω is the pump rotational speed.This normalized pressure-flow dependence (FIG. 3) allows accurate pumppressure drop calculations if the pump flow and speed is known. Thisfunction can be obtained from previous bench pump test data.

The relationship between the motor power and the pump flow rate can beexpressed as follows:Q=φ(S _(elec)/ω²)  (4)where S_(elec) is the average electrical power of the motor in timeinterval T, ω is the pump rotational speed, and φ is a monotonicfunction within the pump operating range. For a given motor-pump systemφ is obtainable from pump test data using, for example, a curve fit.

Flow versus average power relationships can be found experimentally fora variety of rotodynamic blood pumps driven by brushless DC motorshaving any type of phase current waveforms. For a properly designedmotor-pump system this dependence is close to linear as shown in FIG. 4.However a non-linear power-flow dependence can also be derived from pumptest data. This has particular advantages over previously described useof motor current and speed to derive pump flow. The method of flowcalculations described in U.S. Pat. No. 5,888,242 assumes that brushlessDC motor has a sinusoidal back electromotive force (EMF) and sinusoidalphase current waveforms. Most brushless DC motors do not have sinusoidalback EMF and current waveforms. Moreover, when pulse width modulation(PWM) is used to control the motor excitation power the motor phasecurrent may be irregular and prone to spiking (FIG. 5).

The average electrical power consumed by a brushless DC motor is foundby:S _(elec) =v*I  (5)where v is the motor driver bus voltage, and I is the motor average (orDC) current within a certain time interval, T.

The time interval T should be small enough to allow properrepresentation of the pump flow, power and speed pulsatility associatedwith the heart residual function. On the other hand, T should be largeenough to ignore power spikes associated with the motor commutation.This defines the requirements for low-pass filter signal conditioning aswell as analog to digital converter (ADC) sampling rate that is fasterthan 2/T samples/sec. This arrangement provides proper representation ofthe pump performance variation during a cardiac cycle while ignoringpower disturbances associated with motor commutation and excitationpower control by the PWM. In the preferred embodiment the time intervalis120/(NP*ω _(min))<T<30/HR _(max)where NP is number of pump motor magnetic poles; ω_(min) is the minimumpump operating rotational frequency, RPM; HR_(max) is the maximumconsidered patient's heart rate, BPM.

Using motor power instead of motor current (the latter was described inU.S. Pat. No. 5,888,242 by Antaki et al.) for pump flow calculationmeans that the calculation is independent of system voltage variations.Voltage variations that occur while LVAD systems are run on portable orimplantable batteries can significantly change the current required tomaintain the same pump steady state power level. Therefore, thepreferred embodiment has a superior accuracy for flow calculationscompared to methods that use current to calculate pump flows. Thisapproach also eliminates the need for driver voltage regulation that isnecessary if the pump flow rate is calculated using the motor current.

With reference to FIG. 6, a blood pump 10 is driven by a brushless DCmotor 12. In the preferred embodiment, the blood pump 10 is an implantedcentrifugal blood pump. Alternately, mixed or axial flow blood pumps canbe used. A power supply 14 supplies power to the motor 12 via a motorwinding commutation circuit 16. The circuit 16 acts as a power regulatorand is controlled by a microprocessor or microcontroller 18.

The microprocessor 18 receives information from three sensors. A currentsensor 20 provides a current waveform to the microprocessor 18. Thecurrent waveform is the current supplied to the motor 12 by the circuit16 over a period of time. A voltage sensor 22 provides themicroprocessor 18 with a voltage waveform that represents the voltagesupplied to the motor 12 over a period of time. A frequency sensor 24supplies the microprocessor 18 with a frequency waveform that tracks thespeed of the motor 12 over a period of time. Preferably, the frequencysensor 24 is a back electromotive force (EMF) or a rotor position sensorsuch as a Hall effect sensor that calculates a revolutions per minute(RPM) measurement to provide timing for motor winding commutation.Alternately, an optical sensor can be used to track the speed of themotor.

Within a certain time interval (e.g., every ten seconds), themicrocontroller continuously receives a signal from the current sensor20 representing mean DC motor current waveform, a signal from thefrequency sensor 24 representing the pump rotational speed waveform, anda signal from the power supply representing the power supply voltage.

Based on these three inputs, the microcontroller 18 calculates a motorinput power, S_(elec), using equation (5); an instantaneous pump flow,Q, from equation (4); the average pump flow over the time interval; thepatient heart rate HR, by counting the number of pulses in any one ofthe power, current, flow or speed waveforms that are associated with theresidual ventricular function or by finding the waveform fundamentalfrequency; a pump pressure differential, P, using the pump pressure-flowrelationship of equation (3) and a required or target pump flow,Q_(target), for current physiologic conditions according to equation (1)or (2). The microcontroller also analyzes the pump performance andchecks for predetermined patient safety conditions using power, speed,flow and pressure waveforms. Finally the microcontroller implements achange in power delivered to the motor as needed in order to firstcorrect any detected operating conditions or if all conditions are met,to match the required flow, Q_(target).

In case of a software or hardware failure associated with themicrocontroller, the winding commutation circuit 16 providesuninterrupted pump performance at the predetermined (motor power) level.Thus, improved reliability is provided with this arrangement.

In the preferred embodiment, the pump motor input power is the onlydirectly controlled parameter. By varying the pump performance throughthe motor power, a target pump blood flow is maintained based on thederived physiologic HR or HR/P ratio. There are, however, fourfunctional conditions that must be met before implementing any change inmotor power to achieve the target pump flow. These conditions, if notmet, implement a change in motor power to correct the condition and takepriority over motor power changes to achieve target flows.

The first condition is the prevention of ventricular suction. Thisrequires an immediate pumping power reduction if either such conditionor prior to the onset of such situation (a pre-suction condition) isdetected.

The second condition is related to reverse flow through the pump. Aninstantaneous reverse flow is avoided in order to achieve proper flowcirculation through the pump. If instantaneous reverse flow at any timeexceeds a certain limit the controller increases pumping by deliveringmore power to the motor assuming that the first priority condition issatisfied.

A third condition is the microcontroller response to tachycardia orbradycardia or to errors in the HR calculation. If the calculated HRexceeded preset maximum or minimum values, the controller decreasespumping by delivering less power to the pump motor. The maximum andminimum safe HR in the preferred embodiment are 180 and 40 bpmrespectively but can be adjusted to the needs of the individual patient.

The fourth condition is to ensure that pump operational speed ismaintained within a predefined range for ideal system performance. Thelimits may depend on both patient conditions and pump technicalcharacteristics. In the current embodiment, minimum and maximumallowable speed limits were set at 2200 RPM and 3200 RPM, respectively,but can be adjusted for individual patient needs. Any requested changein power delivered to the motor based on target flow requirements or anyof the above priority conditions will not be implemented if they willcause pump speed to exceed these limits.

Prevention of ventricular suction is implemented by detection of apre-suction condition. In the preferred embodiment, pre-suctiondetection is based on the assumption that low pump flow pulsatility isassociated with a completely unloaded ventricle, low intraventricularpressure and steady state (nonpulsatile) systemic pressure. Anysignificant increase in pump flow after depulsing the circulation couldlead to complete or more likely partial collapse of the ventricle wallinto the pump inflow cannula orifice.

The flow pulsatility can be defined as the following:DQ=(Q _(peak(+)) −Q _(mean))/Q _(mean).  (6)where Q_(mean) is the mean flow rate for all cardiac cycles recordedover a given control cycle and Q_(peak(+)) is the average of the maximuminstantaneous pump flow value within each cardiac cycle over the givencontrol cycle. This peak flow is associated with ventricular systolewhen the pressure across the pump is minimal. DQ values below apredetermined limit are used to detect a pre-suction condition andrequire an immediate pumping power reduction.

It is understood that there are other ways to determine the pulsatilityof a waveform using its extreme and mean values as well as time-basedparameters. For example, the following expressions for pump flowpulsatility can also be used to detect a pre-suction conditions:DQ ₁=(Q _(peak(+)) −Q _(peak(−)))/Q _(mean),DQ ₂=(Q _(peak(+)) −Q _(peak(−)))/Q _(peak(+)),DQ ₃=(Q _(mean) −Q _(peak(−)))/Q _(mean), etc.where Q_(peak(−)) is the average of the peak minimum instantaneous flowrates within each cardiac cycle recorded over a given control cycle.This peak minimum flow is associated with ventricular diastole when thepressure across the pump is maximum.

If suction does occur rapidly before an adequate control response basedon low pulsatility limits (DQ), the pump flow will drop significantlydue to inflow cannula occlusion. If the pump flow reduces below apredetermined absolute minimum flow Q_(absmin), complete or significantinflow cannula occlusion is detected and requires an immediate pumpingpower reduction. The value of Q_(absmin) is adjusted to individualpatient physiology.

In the case of partial inflow cannula occlusion, the flow pulsatilitymay remain high, masking DQ limit detection. However, if the pump flowreduces below a predetermined relative minimum flow Q_(rel min) expectedfor the current pump operating speed, partial inflow cannula occlusionis detected and again requires an immediate pumping power reduction. Therelative minimum flow Q_(rel min) expected for any current pumpoperating speed is defined as the following:Q _(rel min) =A*(ω−ω₀)^(n)  (7)In the preferred embodiment, A=1.5, ω₀=1.2 KRPM, n=2, but can bedifferent. If at any given speed within the controller allowed range,the pump flow Q<Q_(rel min) the controller reduces power to the motoruntil the flow is restored to Q>Q_(rel min).

An additional indicator of either a suction or a pre-suction conditionis the absolute value for flow pulsatility defined as the following:absDQ=(Q _(peak(+)) −Q _(peak(−)))where Q_(peak(+)) and Q_(peak(−)) are the average of the peak maximumand peak minimum instantaneous flow rates within each cardiac cyclerecorded over a given control cycle. Any absDQ values below apredetermined limit, can be used to detect such conditions and requirean immediate pumping power reduction.

The preferred embodiment requires three input variables: motor currentwaveform, motor speed waveform and power source (driver) voltage, fromwhich motor power, pump flow, ventricular contraction rate and systemicpressure are determined. This removes the technical and reliabilityproblems associated with direct flow, pressure and heart rate sensing.The motor speed output is available in most brushless DC (BLDC) motordrivers; motor current and power supply voltage sensing are alsoperformed within the driver. This brings all required sensing inside themotor controller/driver and simplifies the system configurationsignificantly. It is to be appreciated that the preferred embodiment isindependent of the motor control mode: the motor speed and current aswell as the driver voltage may vary during a cardiac cycle or remainconstant.

In the preferred embodiment, the current, voltage and speed waveformsare employed to find the heart rate, motor power, mean pump flow andsystemic pressure. Therefore the waveform acquiring interval has to belong enough to determine the heart rate and the average pump flow.Assuming that the normal heart rate is unlikely to be less than 60 BPM,a 6-to-12 second time interval for speed, current and voltage waveformrecording is reasonable. This is also adequate to meet a requiredphysiologic dynamic response of a patient to changing heart rate,preload and afterload conditions.

Because the instantaneous maximum and minimum peak pump flows have to becalculated using the sensed voltage, current and speed waveforms, signalnoise may lead to unacceptable errors which result in improper systemfunctioning. In the preferred embodiment, waveform filtering withlow-pass filters 26 is performed on each of the three waveforms beforebeing digitized by analog to digital converters 28 for analysis by themicroprocessor 18.

Derived from equation 4, the steady state power-flow curves can be usedfor mean, peak, and minimum pump flow calculations. The accuracy of thepump flow calculation using equation (4) is affected by blood viscositychanges. The blood viscosity depends on blood hematocrit level (FIG. 7).As blood viscosity increases, more motor power is required for pumpingat the same blood flow rate. Therefore, for accurate pump flowcalculations a correction factor C_(h) can be added to expression (4):Q=φ(S/ω2)+ChA correction factor such as a 0.1 L/min per each percent of hematocritchange from a baseline is preferred.

Since the instantaneous pump flow is known, the pump inlet-outletpressure difference waveform is calculated using the normalizedpressure-flow curve (3). The estimated mean systemic arterial pressureis derived as the maximum pump inlet-outlet pressure difference, P andis defined as:P=max{P _(norm)*ω^(n)}  (8)This estimate of systemic arterial pressure is used to calculate(M=HR/P) and then the required pump flow Q_(target) is calculated usingequation (2).

As discussed above, a patient heart rate can be obtained by performing aharmonic analysis, a fast Fourier transform or equivalent analysis, onthe motor current waveform. The waveform fundamental frequencycorresponds to the heart rate. The required pump flow Q_(target) iscalculated using equation (1) or (2).

After the mean pump flow (Q_(mean)) is calculated, the flow pulsatility(DQ) as in equation (6) is found. Also the error between the actual pumpflow and the required flow, Q_(target) can be calculated as:ΔQ=(Q _(Target) −Q _(mean))/Q _(mean)  (9)

Since all necessary parameters—Q_(mean), Q_(peak), Q_(min), AbsDQ, DQ,ΔQ, and Q_(target) are calculable, logic analysis is performed and apower control signal is produced for each data acquisition interval ofsix to twelve seconds in the preferred embodiment. With reference toFIG. 8, the logic analysis starts with the calculations 30 of power,heart rate, flow, pressure, pulsatility, minimum and maximum peak flow,and target flow from the received frequency (ω), current (I), andvoltage (V) waveforms 32. The calculated data is then analyzed to see ifany of a set of four priority conditions are not met. First, the speedof the motor is monitored to be sure it is within derived parameters 34,36. If it is greater then the set maximum, ω_(max), then the power isdecreased at step 38 and an alarm may be provided. If the frequency isless than the minimum ω_(min) power is increased at step 40 and, again,an alarm may be provided.

Second, pulsatility DQ is checked to make sure it is not less than apreset limit DQ_(min) indicating onset of inflow cannula suction at step42. If it is, then power is decreased as represented at step 44 and analarm may be triggered. Next, the flow rate is compared to the absoluteand relative minimum flow rates for a given speed 54, 56. IfQ<Q_(abs min) or Q<Q_(rel min), then suction is detected, powerdecreased and an alarm can be actuated 58, 60.

Third, the calculated HR is compared to maximum and minimum limits forarrhythmia or fibrillation detection 62. If HR is out of preset limitsthen power is decreased as represented in step 64 an alarm can beactuated.

Fourth, peak minimum flow rate is compared to a pre-determined maximumallowable reverse pump flow 46. If Q_(peak(−))<Q_(min), then power isincreased as indicated in the flow chart at 48 and an alarm may beactuated.

Finally, if all of the above priority conditions are met, then the powerto the motor can be changed to achieve the target flow based on therelative difference between the actual pump flow and the required flow,ΔQ 50, 52. If any of the above conditions were not met, then anyassociated change in pump power level to correct the condition takespriority over that to achieve the target flow, and the control loopbegins again.

It is to be understood that additional safety and control conditions canalso be utilized in the proposed system. For example, if the motor powerlevel is outside of a predetermined range, an alarm signal is initiatedand the system is switched to a safety mode.

The physiologic control algorithm of the present invention uses a“sensorless” approach, i.e., no pressure or flow sensors, and no ECGsignal is required. The pump speed, motor voltage, and current waveformsare used for heart rate, pump flow, and patient pressure estimates. Thecontroller analyzes this information and develops a control signaltargeting a pump flow that is appropriate to the patient condition. Thispump flow is a predetermined function of the patient pressure and/orheart rate and also depends on the size and degree of heart failure ofthe patient. Several limiting safety conditions and correction factorsare involved in the control signal development.

Several significant advantages are attained relative to knownarrangements.

First, using the pump flow instead of the pump speed as the controlobjective advantageously provides a responsive and flexible controlsystem that can automatically adjust the pump performance according tochanging patient condition (i.e., enhanced left ventricle contractilitydue to the heart recovery). This also eliminates the need for the speedstabilizing circuitry (speed control loop) in the motor drive since thespeed may vary during the cardiac cycle.

Second, the developed algorithm is applicable to a wide variety ofbrushless motors driven by currents having waveforms of any shape,rather than being limited to a sine-wave current type of motor. This isadvantageous since most brushless motors are not of this type.

Third, the present control algorithm avoids the potentially dangeroussituation of ventricular collapse by detecting conditions which precedethat event, as opposed to incrementing pump speed until such a situationis detected as suggested by others.

Fourth, the control methodology of the present invention also simplifiesthe control system by eliminating the need for speed stabilizingcircuitry and voltage regulating circuitry.

Fifth, in addition, the control method disclosed herein is independentof voltage variations that can be significant when the pump is poweredby batteries.

Sixth, still further, the present control arrangement is deemed moreaccurate and consequently safer for patients since it recognizes andtakes into account patient blood viscosity variations and is alsoindependent of blood flow inertia.

The invention has been described with reference to the preferredembodiment. Modifications and alterations will occur to others upon areading and understanding of the preceding detailed description. It isintended that the invention be construed as including all suchmodifications and alterations insofar as they come within the scope ofthe appended claims or the equivalents thereof.

1. A cardiac assist apparatus comprising: a blood pump; a drive unitthat drives the blood pump; a power supply that supplies power to thedrive unit; a frequency sensor that senses a rotational speed of theblood pump; a current sensor that senses an average direct currentwaveform of the drive unit; a power supply voltage sensor that senses apower supply voltage; and a blood pump controller that; (i) receivesdata from the frequency sensor, current sensor, and the power supplyvoltage sensor; (ii) passively determines, from said data, a physiologicstatus of a patient without affecting one or more of the following: thepump's performance and the patient's physiology, the patient'sphysiologic status comprising one or more of the following: a drive unitinput power for determination of pump flow rate; a heart rate; a pumpflow rate; and, a pressure differential; and (iii) calculates a targetflow based on the physiologic status; and (iv) automatically adjusts theblood pump to the target flow.
 2. The cardiac assist apparatus as setforth in claim 1, wherein the controller estimates a drive unit inputpower, a flow rate, a patient heart rate, pressure and the target flowrate.
 3. The cardiac assist apparatus as set forth in claim 1, furthercomprising a drive unit winding commutation circuit for relaying thetarget flow to the drive unit motor.
 4. The cardiac assist apparatus asset forth in claim 1, wherein the drive unit includes a brushless DCmotor.
 5. The cardiac assist apparatus as set forth in claim 1, whereinthe power supply includes at least one rechargeable battery.
 6. Thecardiac assist apparatus as set forth in claim 1, wherein the frequencysensor includes one of a back EMF, Hall and optical sensor.
 7. Thecardiac assist apparatus as set forth in claim 1, wherein the patient'sphysiological status is a heart rate, and the blood pump controllercalculates the target pump flow rate based on the determined heart rate.8. The cardiac assist apparatus as set forth in claim 7, wherein thetarget flow rate is within a pre-determined margin.
 9. The cardiacassist apparatus as set forth in claim 8, wherein the margin comprisesflow rates corresponding to preset maximum and minimum heartbeats perminute.
 10. A method of controlling blood flow with a blood pumpcomprising the steps of: sensing a current waveform of a drive motor ofthe blood pump; sensing an input voltage waveform to the drive motor;sensing a rotational frequency waveform of the drive motor; passivelydetermining, at a blood pump controller, a physiologic status of apatient based on one or more of the sensed current waveform, inputvoltage waveform, and rotational frequency waveform, the patient'sphysiologic status comprising one or more of the following: a drivemotor input power of pump flow rate; a heart; a pump flow rate; and, apressure differential wherein said step of determining is performedwithout affecting one or more of the following: the pump's performanceand the patient's physiology; and calculating a target flow based on thephysiologic status; and automatically adjusting the blood pump to thetarget flow.
 11. The method as set forth in claim 10, further comprisingthe step of analyzing the pump's performance to verify that the pump isoperating within predetermined conditions.
 12. The method as set forthin claim 10, wherein the predetermined conditions comprise one or moreof the following: a minimum flow rate; a maximum flow rate; a minimumpulsatility; a minimum pump rotational speed; and, a maximum pumprotational speed.
 13. The method as set forth in claim 12 wherein thepredetermined conditions comprise one or more of the following: absoluteminimum flow rate, and relative minimum flow rate.
 14. The method as setforth in claim 13 further comprising the step of calculating therelative minimum flow rate as Q_(rel min)=A*(ω−ω₀)n wherein A, n and ω₀are constants and ω is the pump rotational speed.
 15. The method as setin claim 13, further comprising the step of reverting to a pre-selecteddefault power level decrease in response to detecting ventricularsuction.
 16. The method as set forth in claim 10, wherein the motorinput power is calculated using the equation:S _(elec) =v*I where S_(elec)=v*I is an average electrical power of themotor in a time interval T, v is the measured voltage, and I is anaverage current over the time interval I.
 17. The method as set forth inclaim 16, wherein the patient's physiologic status is a heart rate, thestep of calculating the heart rate comprises analyzing one or more ofthe following for a fundamental frequency: the current waveform, inputpower, rotational frequency, pressure differential and pump flowwaveform.
 18. The method as set forth in claim 10, wherein the pumppressure differential P is calculated using the equation:P _(norm)=ψ(Q _(norm)) where normalized pump pressure Pnorm=P/(σ*ω2),normalized flow rate Q_(norm)=Q/ω, and ω is the pump rotational speed, σis the blood density.
 19. The method as set forth in claim 10, whereinthe target pump flow is calculated using one of the following equations:Q _(target) =C ₁ *HR+C ₂; and,Q _(target) =f(M); where Q_(target) is the desired flow rate, C₁ and C₂are constants determined by patient heart condition and size, M=HR/Pwhere P is the pressure and HR is the heart rate.
 20. The method as setforth in claim 10, wherein the step of calculating the target pump flowfurther comprises the step of tracking the patient's heart rate andbasing the flow rate on the heart rate.
 21. The method as set forth inclaim 10, further comprising the step of reverting to a pre-selecteddefault pump rotational speed in the event of a controller failure. 22.The method as set forth in claim 10, further comprising the step ofreverting to a pre-selected default motor input power in the event of acontroller failure.
 23. The method as set forth in claim 10, furthercomprising the step of reverting to a pre-selected default PWM dutycycle in the event of a controller failure.
 24. The method as set forthin claim 10, wherein the pump flow is calculated using the equation:Q=φ(S/ω ²) where S=V*I, S is average electrical power, V is motor drivenbus voltage, I is motor average current within a certain time interval,Q is pump flow, and ω is the pump rotational speed.